Radiation therapy, which is one of three main methods of treating cancer, together with surgery and chemotherapy, is currently carried out predominantly with high energy x-rays of one to several MeV energy produced by special x-ray generators employing electron linear accelerators (“linacs”) of several MV high voltage. MeV x-rays have good attributes for use in radiation therapy, in particular, high tissue penetration and a robust sparing of the first few millimeters of shallow tissues, generally known as a “skin-sparing effect.” They also have several shortcomings, most significantly, the normal, non-targeted tissue that is located proximal, distal, and lateral to the target receive excessive radiation damage as described further herein.
This is because the mode of interaction of the high energy x-rays that are produced, typically 1-4 MeV, is Compton scattering and not photoelectric. As a result, the dose distribution produced in a patient's body is mostly from multiple Compton scattering from a wide range of angles and, therefore, is not well-confined within the target.
In particular, the doses produced at the target tissue by MV sources do not sharply fall at the target's edge. Instead, the dose distribution at the target's edge is rather blunt-edged. Quantitatively, the so-called “80%-to-20% dose falloff” produced at the target by high energy x-rays is typically 2-5 mm. In addition, the beam-shaping collimators, so-called “multi-leaf collimators,” required to produce the high-energy beam profiles, consist of heavy, thick “leaves” which do not lend themselves to production of fine exposure profiles. Because these collimators fail to produce beam-exposure profiles with fine contours, unnecessary radiation dose is delivered to normal tissues, especially when small targets are exposed. Such large falloffs result in unnecessary and undesirable dose being delivered to the tissues located in the immediate neighborhood of the target.
Further, because high energy x-rays have little preferential absorption in heavier elements compared to the light elements that constitute most of the tissues, the concept of tumor-dose enhancement by the introduction of contrast agents to the tumor such as iodine and gold cannot be effectively implemented when the radiation type is high energy x-rays. In addition, although the large penetration of the dose from high-energy x-rays to tissue depths is considered an advantage for thick targets, for thin tumors the shallow dose falloff of the high energy x-rays with depth is a negative effect, allowing the exposure to high radiation dose of all tissues positioned distal to the target. FIG. 1 illustrates dose penetration 10 in tissues for different high-energy MeV x-ray beams 12, compared to the dose penetration curve for an orthovoltage tube 14.
Before MV x-ray machines were developed (around the mid-20th century), x-ray generators of lower energy, called “orthovoltage” x-ray machines or tubes were used for radiation therapy. The acceleration voltage of these early x-ray machines was rather small, mostly up to 250 kVp, producing x-rays with a median energy, or mean energy, of about 110 keV. These beam energies were too low to penetrate deep in the tissue, and also lacked the beam sparing effect of the shallow tissues that the high-energy MV x-rays exhibit, in fact lower than that shown in FIG. 1 for orthovoltage x-rays. As a result, the skin and the normal tissues proximal to the target received significant radiation damage. FIG. 1 compares the dose penetration in tissues from high energy x-rays produced by electron linacs to that from a 300 kVp orthovoltage tube filtered moderately, labeled by half-value layer (HVL) in copper as “3.0 mm Cu HVL.”
To address the damage to healthy skin tissue using orthovoltage x-rays, a so-called “grid therapy” was developed. Conventional grid therapy used a metal or lead grid with openings of at least 1.0-1.5 cm diameter to ameliorate the skin damage that occurred in treating deep tumors. However, the orthovoltage grid therapy techniques offered little, if any, tissue-sparing to healthy subcutaneous tissue proximal to the target, and thus did not solve the problem of damage to the normal tissues proximal to deep tumors. Furthermore, no method or system was contemplated for controlling the tissue depth at which a therapeutic dose could be produced across a target by the merging of the beams exiting the grid.
Accordingly, there is a need for a method and system for performing radiotherapy using orthovoltage x-rays for effectively treating tumors while sparing both the skin and tissue proximal to the target. There is also a need for a system and method for controlling the tissue depth at which a therapeutic dose of orthovoltage x-ray radiation can be delivered to the target while sparing tissue proximal to the target. The development of such improved orthovoltage x-ray systems may provide not only benefit to a wide range of clinical applications by reducing dose to the non-targeted tissues, but also a low-cost and compact solution for performing radiotherapy to effectively treat tumors, as well as neurological targets.